1. Field of the Invention
The present invention relates to totally implantable hearing devices generally and particularly to totally implantable cochlear implants which stimulate the ossicular chain of the middle ear for restoring partial hearing loss and for totally implantable devices which stimulate the cochlear nerve endings to restore total hearing loss.
2. Description of the Prior Art
Conventional prior art hearing aids are composed of a microphone, an amplifier, a battery as a power source, and a speaker or earphone (commonly referred to as a receiver in the hearing aid industry). Known implantable hearing device have the same basic components, except that the speaker is replaced by a driving vibrating component, such as an electromagnetic coil or a piezoelectric system of biomorph design. Environmental sound energy, as it passes through either device, is converted by the microphone into an electrical signal which is routed to an amplifier. In the conventional hearing aid, the speaker converts the amplified electrical signals into acoustic energy, which is then transmitted to the tympanic membrane and ossicular the speaker is eliminated, being replaced by the vibratory component which drives the ossicular chain.
Partially implantable middle ear hearing devices are also known. These devices have a small, lightweight high coercivity magnet effectively glued to the ossicular chain by a bio compatible bonding material such as METABOND or SUPERBOND adhesives manufactured in the USA and Japan respectively. The magnet is driven by an air core electromagnetic coil optimally spaced from the target magnet at a distance of approximately 1 mm. There is no contact between the air core coil and the target magnet.
These devices have an external unit and an internal unit. The external unit receives, amplifies, and transmits sound energy as radio frequency signals. The external unit consists of a microphone, a radio frequency (RF) amplifier, a transmitting antenna, and a battery. Using existing microchip technology, these components are miniaturized to a unit with dimensions of 10.times.10.times.5 mm without the battery.
The internal unit consists of a receiving antenna, a titanium support, implanted electronics, an electromagnetic (EM) transducer (driving coil), and the high coercivity magnet. The electronics (diode and capacitor), driving coil, and magnet are hermetically sealed in a helium filled laser-welded titanium case. A glass-insulated feed-through attaches the electronics to silicone or polytetrafluoroethylene-coated platinum iridium or stainless steel wires of the receiving antenna. The precise alignment of the transmitting (external) and receiving (internal antennae permits transcutaneous transfer of the sigma delta modulated radio frequency signal (8 to 10 .sub.MHZ). The implanted electronics function to receive the radio frequency signal that has been processed by the external electronics and to transform this energy into an audio frequency input to the driving coil. The driving coil in turn creates a magnetic field, which activates the target magnet attached to the body of the incus. Through the ossicular chain, the vibrations are transmitted to the inner ear fluids, activating the organ of Corti.
The magnet used is a neodymium-iron-boron (NdFeB) permanent magnet of great coercive force and high flux density. The magnet, weighs 8.0 mg, is hermetically sealed in a laser-welded 6-mg titanium case containing a helium atmosphere. Two of such magnets are used, stacked on top of each other. On the basis of fresh human cadaver studies, the magnet-titanium assembly weight load of 65 mg and 110 mg has a negligible effect at the malleus and incus, respectively, on the frequency response.
The external electronics associated with the transducer are designed to apply only push forces on the magnet-incus assembly. An air-core coil placed in the attic of the middle ear is used because it does not exert a constant bias force on the ossicular chain. If the system is idling, there is no steady force applied to the incus-magnet assembly. In order to determine the size of the driving coil, 20 preserved human cadaver temporal bones were microsurgically dissected. Measurements were made of the mastoid cavity, antrum, attic, and body of the incus. Owing to the anatomic characteristics of the attic, the outside diameter of the driving coil assembled in a titanium case was limited to 5.0 to 6.0 mm. outside diameter. Initially, an efficient coil was built with a 3.0-mm outside diameter, 0.75-mm inside diameter, and a length of 1.0 mm, composed of 2200 turns of 52 AWG copper wire with 600 ohms resistance. A more efficient coil with 2668 turns of 52 AWG copper shire with 875 ohms resistant was later built for short-and long-term animal experimentation. Computer simulations were instrumental in the selection of this coil design. For the human device a coil with 3800 turns and a resistance of 1415 ohms is preferably used. This coil has been tested experimentally and found to generate 76% more force when compared to the 2668 turn coil.
These described devices while operating to provide restoration of partial hearing loss did not have any means to activate the cochlear nerve endings in patients who had total hearing loss due to a malfunctioning cochlea. Also, these devices had no provision for adjusting the device to accommodate growth of the patient over the years which is a problem especially prevalent in implants for growing children. Further, these devices were sensitive to magnetic resonance imaging (MRI)interference and any patient with such devices when subjected to an MRI came out of the scan with an inactive hearing device.